Digital radiography (DR) is increasingly accepted as an alternative to film-based imaging technologies that rely on photosensitive film media. With DR, a detector panel or Flat Panel Detector (FPD) provides an array of sensing circuits that convert levels of radiation exposure captured on radiation-sensitive layers to electronic image data. The image photosensors are typically metal-insulator-semiconductor (MIS) diodes or PIN (P-type, Intrinsic, N-type) diodes or some other photosensor type. The array of image pixel data that is obtained from the DR detector is then stored in memory circuitry for subsequent read-out, processing, and display on suitable electronic image display devices.
FIG. 1 illustrates a cross-sectional view of a conventional DR panel 10 with a flat panel detector 20. A scintillator 14 has a material, such as gadolinium oxisulfide, Gd2O2S:Tb (GOS) or cesium iodide, that absorbs x-rays incident thereto and converts the x-ray energy to visible light photons. Flat panel detector 20 is physically adjacent to the scintillator (GOS) layer, and includes an array of light sensitive pixels 24 arranged in a matrix of rows and columns. The pixels 24 are connected to readout elements 25. As shown at enlarged section E of FIG. 1, each pixel 24 has one or more photosensors 22, such as a PIN diode or other light-sensitive component, and an associated switch element 26 of some type, such as a thin film transistor (TFT).
Flat panel detector 20 is typically formed using hydrogenated amorphous silicon (a-Si:H). Pixels 24 on this a-Si:H array record the intensity of the light output from the GOS or other scintillator 14 layer upon absorption of an x-ray. The light sensitive components of the a-Si:H pixels convert the incident light into electrical charge which is stored in the internal capacitance of pixel 24. The magnitude of the stored electrical charge is related to the intensity of the excited light, which is, in turn, related to the intensity of the incident x-rays. The readout of the image information is performed by peripheral electronic circuitry that connects to the edge of the a-Si:H array and is represented as readout elements 25 in FIG. 1. The charge readout from each individual pixel is converted to a digital value in an analog to digital converter (ADC) that is linked to the readout circuitry. The digital values are then transferred to the external system via standard data communication means, such as a wired or wireless data link.
The conventional DR panel receives its operating power from an external source, routed to the panel by means of a cable, tether, or other interconnection device. In the conventional arrangement, operating power is provided continuously to the DR panel. This mode of operation is suitable for conventional, large-scale digital radiographic installations, where the FPD is permanently installed at a predetermined optimum fixed location for patient imaging. This type of installation is typically set up for obtaining a standardized set of radiological images that are routinely needed for a large number of patients. After a warm-up period following power-up, the DR panel reaches a stable operating state that is maintained during and between imaging sessions for a succession of patients. Even when the panel is not capturing images it goes through repeated cycles of reset/refresh, integration, and readout functions, in an automated sequence. Most panels also perform automatic periodic dark calibrations to compensate for temperature drifts and other factors that can affect image quality.
It is known in the art that even continuously operating FPDs exhibit pixel-to-pixel variations in sensitivity and dark current. These variations, sometimes referred to as fixed pattern noise, may result in perceptible nonuniformities in diagnostic X-ray images and thereby interfere with the detection of disease features. Thus, compensation algorithms are necessary, such as those described by James A. Seibert, John M. Boone, and Karen K. Lindfors in “Flat-field correction technique for digital detectors,” Proc. SPIE Vol. 3336, 1998, p. 348-354; by Jean-Pierre Moy and B. Bosset in “How does real offset and gain correction affect the DQE in images from x-ray flat detectors?” Proc. SPIE, 3659, 1999, pp. 90-97; and by Hans-Aloys Wischmann, Hans A. Luijendijk, Henk J. Meulenbrugge, Michael Overdick, Ralf Schmidt, and Kourosh Kiani in “Correction of amplifier nonlinearity, offset, gain, temporal artifacts, and defects for flat-panel digital imaging devices,” Proc. SPIE Vol. 4682, 2002, p. 427-437.
The most basic calibration and correction algorithms generally include 2 steps. First, the dark signal of the detector (that is, the signal in the absence of any X-ray exposure) is obtained. Pixel by pixel variations in the dark signal of the detector are characterized to form a dark or offset map containing the dark variations. The offset map is then subtracted from the X-ray exposure in a process termed dark or offset correction. Second, the variations in the sensitivity of the pixels are characterized. This is done by capturing one or more flat field exposures, which are then offset-corrected. The resulting image is the gain map. In the gain correction step, the offset-corrected X-ray exposure is divided by the gain map. Ideally this two-step procedure compensates for any fixed pattern noise introduced by the detector.
While these two operations seem straightforward, both gain and offset maps have some inherent noise, both may drift over time, and they may exhibit differences depending on previous exposures taken by the detector. Some of these variations are related to the temperature sensitivity of amorphous silicon and to the tendency of this material to trap charge from previous exposures (See, for example, Street, Robert A., Technology and applications of amorphous silicon, Berlin: Springer Verlag; 1999, Chapter 4). Likewise, drift may occur due to readout electronics. Consequently, significant effort has been dedicated to improving the performance and efficiency of the gain and offset corrections.
For example, it is well known in the art that individual dark captures and flat field exposures contain electronic and X-ray quantum noise, respectively. Thus, several captures of each must be averaged to obtain gain and offset maps with reduced noise levels. Noise inherent in those correction maps would propagate to the final corrected X-ray exposure and could potentially interfere with clinical diagnoses. The need for averaging was anticipated by Moy and Bosset; Pieter G. Roos et al., “Multiple-gain-ranging readout method to extend the dynamic range of amorphous silicon flat-panel imagers,” Proc. of SPIE, 5368, 2004, pp. 139-149; and by Tadeo Endo in “Radiological imaging apparatus and method,” U.S. Pat. No. 7,113,565 B2.)
The block diagrams of FIGS. 2 and 3 show conventional approaches for performing offset corrections. Using the sequence shown in FIG. 2, a number n of dark images D are obtained after the actual exposure E. Dark images D are then averaged using the calculation shown, and the average is subtracted from the exposure E data to obtain the offset-corrected exposure image. In the sequence of FIG. 3, n dark images D are obtained prior to the exposure E, and the same combination logic is used to obtain the offset-corrected exposure image.
Some effort has been made to capture the minimum necessary number of flat field and dark images for gain and offset corrections without negatively affecting the noise in the corrected image. Such solutions include frequency decomposition to reduce high frequency noise in the gain map (Brian G. Rodricks, Denny L. Lee, Michael G. Hoffberg, and Cornell L. Williams, “Filtered gain calibration and its effect on DQE and image quality in digital imaging systems,” Proc. SPIE Vol. 3977, p. 476-485) and periodic weighted updates of the existing offset map, as described in U.S. Patent Application Publication No. US2003/0223539 entitled “Method and apparatus for acquiring and storing multiple offset corrections for amorphous silicon flat panel detector,” by Granfors et al. The latter method, wherein a single dark image is captured periodically between exposures and weighted with the existing offset map, is well suited for conventional FPDs running continuously in a stable environment. In this environment, the method described in the '539 Granfors et al. publication captures long term drifts, while reducing noise by averaging multiple dark captures.
Subtraction of the appropriate dark signal in the offset correction is important because any discrepancy between the actual dark level that was present during the exposure and the subtracted offset map is amplified by subsequent correction steps. One of the mechanisms that may change the dark level for a continuously running FPD is image lag, a problem familiar to those skilled in the diagnostic imaging field. Image lag is unwanted charge retention from frame to frame due to incomplete readout of the photodiode, afterglow of the scintillator, trapped charge in the a-Si photodiode and/or other causes. Image lag may be of some concern in cases where dark images are obtained after image capture. The residual image decays over time in a predictable fashion and can be corrected as disclosed by Partain et al. in U.S. Pat. No. 7,208,717 entitled “Method and apparatus for correcting excess signals in an imaging system.” Image lag is proportional to exposure, and its magnitude can be estimated by taking the difference of two dark frames captured at known time intervals after the exposure. Lag correction is mainly of concern for panels running continuously in fluoroscopic mode and panels that switch between high-dose radiographic images and low-dose fluoroscopic images.
While gain and offset corrections pose some challenges for continuously running FPDs, correction algorithms are expected to become more complex as portable, untethered DR panels, which encounter less stable operating conditions, become more prevalent. Untethered DR operation offers some promise of improved patient care, with advantages including improved operator workflow and equipment adaptability. In untethered operation, a portable FPD can be readily positioned behind the patient, rather than requiring the patient to take an awkward position for imaging. In many cases, an untethered flat panel detector can replace the need for multiple conventional detectors, since the same detector can be used both in a wall-mount position and a horizontal table position. The portable, battery-powered FPD has the flexibility of being easily and quickly movable to any suitable location for DR imaging, yet still provides immediate access to the acquired x-ray image. The portable, cassette-type FPD, in turn, allows smaller and more portable x-ray imaging systems to be used. In some cases, portable DR panels can be used where conventional tethered DR panels are not well-suited for patient imaging, and can obviate the need to return to the use of older technologies, such as the use of a storage phosphor computed radiography (CR) X-ray cassette.
Battery power offers considerable benefit, however, there are drawbacks associated with battery use, including the need for battery power conservation when not in use. Battery conservation means that some type of “standby” power level be provided, so that the DR panel can be maintained in a state of readiness, but without drawing the full amount of battery current that is needed for operation until necessary. Any type of power mode switching, however, can have a negative impact on image quality.
Because of the temperature sensitivity of amorphous silicon mentioned previously (cf. R. A. Street reference), the change of power modes, such as to provide “standby” and operation modes, brings with it the likelihood of rapidly changing temperature profiles over the full detector area. This includes both global and local changes, because some electronic components heat up faster than others. Rapid local or global changes in temperature are likely to cause a range of imaging anomalies. DR panel imaging characteristics immediately following a change in operating power can differ measurably from imaging characteristics a few minutes later. This is one reason why the straightforward correction sequences of FIGS. 2 and 3 fall short of what is needed for dark correction in portable DR applications. First of all, workflow considerations may limit the number of dark images that can be taken before the exposure. Ideally, the detector must be ready for the X-ray capture as soon as possible after its transition from the standby mode. Moreover, where there is a rapidly changing temperature profile, a dark image taken immediately after the exposure may not be representative of the dark level that was actually present during the exposure.
The task of properly characterizing and compensating for variations in imaging performance for a portable battery-operated DR panel is complicated by the nature of its use and operation. The panel may be used in different rooms and for different tasks that vary in usage pattern and temperature environment. In an intensive-care unit (ICU), for example, there may be no standard usage pattern or regular timing that could help to predict the amount of compensation needed at any particular point. Instead, use of the DR panel can be more randomized and asynchronous, requiring some adaptive method for proper characterization and calibration.
Thus, although portable, battery-operated DR panels offer clear advantages for operator workflow and improved patient care, these devices present a particular challenge to the task of obtaining a quality diagnostic image. The new set of problems introduced by using on-board battery power for the DR panel requires solutions that minimize the impact of power cycling and uneven heat build-up on the image data that is obtained.